Differential phase-contrast imaging represents an imaging method which has attracted much attention for some time, especially in the Talbot-Lau interferometer arrangement. Thus for example the publication by F. Pfeiffer et al. [1], “Hard X-ray dark-field imaging using a grating interferometer”, Nature Materials 7, 2008, Pages 134 to 137, describes how, with the aid of an interferometric structure, which includes a conventional x-ray tube, three gratings and an x-ray detector, both absorption contrast, differential phase contrast and also dark-field contrast can be reconstructed from the same dataset. Similar information can be found in Joseph J. Zambelli, et al. [2], “Radiation dose efficiency comparison between differential phase contrast CT and conventional absorption CT”, Med. Phys. 37 (2010), Pages 2473 to 2479.
The wave nature of particles such as x-ray quanta allows the description of phenomena such as refraction and reflection with the aid of the complex refraction indexn=1−δ+iβ. 
In this case the imaginary part β describes the absorption on which current clinical x-ray imaging, such as computed tomography, angiography, radiography, fluoroscopy or mammography is based, and the real part δ describes the phase offset, which is observed in differential phase imaging.
An x-ray recording system is known from DE 10 2010 018 715 A1 in which an x-ray recording system for phase-contrast imaging of an object under examination is used for high-quality x-ray imaging, having at least one x-ray emitter with a plurality of field emission x-ray sources for emitting a coherent x-ray radiation, an x-ray image detector, a diffraction grating G1 disposed between the object under examination and the x-ray image detector and a further grating G2, which is disposed between the diffraction grating G1 and the x-ray image detector.
An x-ray recording system, which allows a differential phase-contrast imaging of the type stated at the start to be carried out, is known for example from U.S. Pat. No. 7,500,784 B2, which is explained with reference to FIG. 1.
FIG. 1 shows the typical main features of an x-ray recording system 1 for an interventional suite with a C-arm 2 held by a stand in the form of a six-axis industrial or articulated-arm robot, to the ends of which an x-ray radiation source, for example an x-ray emitter 3 with x-ray tube and collimator, and an x-ray image detector 4 as x-ray recording unit are attached.
By way of the articulated-arm robot known for example from U.S. Pat. No. 7,500,784 B2, which preferably has six axes of rotation and thereby six degrees of freedom, the C-arm 2 can be adjusted spatially in any given manner, for example by being rotated around a center of rotation between the x-ray emitter 3 and the x-ray image detector 4. The inventive x-ray recording system 1 is especially able to be rotated around centers of rotation and axes of rotation in the C-arm plane of the x-ray image detector 4, preferably around the center point of the x-ray image detector 4 and around the center point of axes of rotation intersecting with the x-ray image detector 4.
The known articulated-arm robot has a base stand, which is rigidly mounted on a floor for example. A carousel, able to be rotated around an axis of rotation, is fastened thereto. A robot rocker is attached to the carousel, pivotable around a second axis of rotation, to which a robot arm pivotable around a third axis of rotation is fastened. Attached to the end of the robot arm, pivotable around a forth axis of rotation is a robot hand. The robot hand has a fastening element for the C-arm 2, which is able to be pivoted around a fifth axis of rotation and able to be rotated around a sixth axis of rotation running at right angles thereto.
The realization of an x-ray recording system is not dependent on the industrial robots. Normal C-arm devices can also be used.
The x-ray image detector 4 can be a rectangular or square, flat x-ray detector, which preferably includes a scintillator (e.g. CsI) and an active matrix of photodiodes made of amorphous Silicon (a-Si). However CMOS-based integrating detectors or also counting detectors (e.g. CdTe or CZT and ASIC) can also be used.
Located on a table plate 5 of a patient support table, in the beam path of the x-ray emitter 3, as the object under examination, is a patient 6 to be examined. Connected to the x-ray recording system 1 is a system control unit 7 with an imaging system 8, which receives the image signals of the x-ray image detector 4 and processes them (operating elements are not shown for example). The x-ray images can then be viewed on displays of a monitor array 9. The monitor array 9 can be held by way of a ceiling-mounted, longitudinally-movable, pivot, rotation and height-adjustable support system 10 with outriggers and a lowerable support arm.
Instead of the x-ray recording system 1 with the stand in the form of the six-axis industrial or articulated-arm robot as shown by way of example in FIG. 1 the x-ray recording system 1 can, as shown in FIG. 2 in simplified form also have a normal ceiling-mounted or floor-mounted holder for the C-arm 2.
Instead of the C-arm 2 shown by way of example, the x-ray recording system 1 can also have separate ceiling-mounted and/or floor-mounted holders for the x-ray emitter 3 and the x-ray image detector 4, which are electronically rigidly coupled for example.
In the current arrangements focused on for clinical phase-contrast imaging conventional x-ray tubes, currently available x-ray image detectors, such as are described for a example by Martin Spahn [3] in “Flat detectors and their clinical applications”, European Radiology, Volume 15 (2005), Pages 1934 to 1947, and three gratings G0, G1 and G2 are used, as is explained below with reference to FIG. 2, which shows a schematic structure of a Talbot-Lau interferometer for differential phase-contrast imaging with the extended tube focus, gratings G0, G1 and G2 and pixelated x-ray image detector.
The x-ray beams 12 emerging from a tube focus 11 of the non-coherent x-ray emitter 3, for generation of coherent radiation, penetrate an absorption grating 13 (G0), which brings about the local coherence of the x-ray emitter 3, and also an object under examination 14, for example the patient 6. Through the object under examination 14 the wave front of the x-ray beams 12 is deflected by phase offsetting in such a way as illustrated by the normal 15 of the wave front without phase offsetting, i.e. without object, and the normal 16 of the wave front with phase offsetting. Subsequently the phase-offset wavefront passes through a diffraction or phase grating 17 (G1) with a grating constant adapted to the typical energy of the x-ray spectrum for generation of interference lines or an interference pattern 18 and in its turn an absorbing analyzer grating 19 (G2) for reading out the interference pattern 18 generated. Different interference patterns 18 arise with and without object. The grating constant of the absorbing analyzer grating 19 is adapted to that of the phase grating 17 and the remaining geometry of the arrangement. The absorbing analyzer grating 19 is disposed for example in the first or nth Talbot spacing (order). The absorbing analyzer grating 19 in this case converts the interference pattern 18 into an intensity pattern, which can be measured by the x-ray image detector 4. Typical grating constants for clinical applications lie at a few μm, as can also be found in the cited literature references [1, 2].
If the x-ray emitter 3 is sufficiently coherent, i.e. the tube focus 11 of the x-ray emitter 3 is sufficiently small and the generated radiation power is still sufficiently large, it is possible to dispense with the first grating G0, the absorption grating 13.
The differential phase offsetting is now determined for each pixel of the x-ray image detector 4 according to the prior art by, through what is referred to as “phase stepping” 20, which is indicated by an arrow, the absorbing analyzer grating 19 (G2), being displaced in a number of steps (k=1, K, with e.g. K=4 to 8) by a corresponding fraction of the grating constant at right angles to the radiation direction of the x-ray beams 12 and lateral to the arrangement of the grating structure and by, for this configuration, the signal Sk arising during the imaging in the pixel of the x-ray image detector 4 being measured and the interference pattern 18 produced being sampled. For each pixel the parameters of a function describing the modulation (e.g. sine) are then determined by a suitable fit method, an adaptation or compensation method, to the signals Sk thus measured. These parameters are usually the amplitude A, the phase Φ and the average intensity I.
The comparison of specific derived variables from these fit parameters for each pixel, once with and once without object (or patient), then enables three different images to be created:
Absorption image,
Differential phase-contrast (DPC) image and
Dark-field image.
The visibility, i.e. the standardized difference between the maximum and minimum signal (or more precisely: amplitude standardized to the average signal) is in this case a measure for characterizing the quality of a Talbot-Lau interferometer. It is defined as the contrast of the sampled modulation
  V  =                              I          max                -                  I          min                                      I          max                +                  I          min                      =          A              I        _            
References in this document to an image may possibly mean the triumvirate of absorption image, DPC image and dark field image.
The realization of the method presents many challenges, but in particular has two decisive disadvantages:                The absorbing analyzer grating G2 must be moved into different positions and then an x-ray acquisition carried out in each position. Such a method is thus conceivably unsuitable for moving objects (such as non-anaesthetized patients or patient organs, e.g. heart, lungs) if the object moves between the different measurements even by small distances. Likewise such an arrangement is unsuitable on account of the mechanical movement of the absorbing analyzer grating G2, in order to make possible real-time imaging or imaging at higher image frequencies of for example 15 images per second (I/s) or also 60 to 100 I/s. 3D imaging, in which x-ray tube and x-ray detector are rotated continuously around the patient, are also not possible in this way.        The fact that the absorbing analyzer grating G2 is an analyzer grating with areas in which it is transparent for x-rays and in other areas is as non-transparent as possible means that dose (typically 50%) is lost behind the object or the patient, which is not effective in the image.        